Laser-ultrasonics is a well established technique first developed for the non-destructive testing of industrial materials like metals, plastics and polymer-matrix composite materials. A typical laser-ultrasonic system is composed of two lasers and a phase demodulator, as sketched in FIG. 1. In laser-ultrasonics, ultrasound is generated and detected at a distance. The ultrasound is usually generated on the surface of the material and the detection of ultrasound is performed on the same surface with a detection laser and a phase demodulator collecting the reflected and/or scattered light. Laser-ultrasonics can be applied to parts having complex shapes without any surface preparation and at high temperature. This technique has been shown to be applicable to real industrial conditions. In particular, the technology has been commercially developed for measuring on-line the wall thickness of steel tubes at 1000° C. moving at 4 m/s and for inspecting polymer-matrix composite materials used in aerospace.
The generation laser emits short optical pulses (from about 100 ns to 100 fs) to produce an ultrasonic wave at the surface of the sample by pure thermoelastic generation, or by thermoelastic generation coupled with some ablation of the surface material. The detection laser then illuminates the surface with a single frequency (or narrow linewidth), moderate power, very stable, long pulse laser. The laser pulse has a pulse duration between 50 and 200 microseconds. The peak power of the detection laser can exceed 1 kilowatt. The detection laser is usually based on a master oscillator power amplifier (MOPA) design which includes a very stable single frequency continuous wave (CW) laser oscillator followed by one or more optical amplification stages. Optical isolators are required between the master oscillator and the amplifying stages.
Light retroreflected or backscattered from the sample is then collected with a large étendue optical system and coupled to an optical phase demodulator. The wavefront insensitive demodulator is usually a confocal Fabry-Perot interferometer (CFPI) or a photorefractive interferometer (PRI). These phase demodulators are able to process beams with large étendue, which is an essential requirement in real applications where the beam reflected from the sample is strongly scattered by rough surfaces. The output signal from the demodulator is then digitally sampled by an acquisition card and processed by a personal computer.
The successful industrial applications reported above, as well as others, were made possible by detection schemes and technology that provide sufficient sensitivity. Adequate sensitivity is obtained by using a demodulator, such as a CFPI or a PRI, that has a large étendue and allows processing a collected beam with speckles. It requires receiving a significant amount of light from the surface, which is realized by having a collecting aperture as large as is practical (which means a large solid angle) and using a relatively high power detection laser. Having sufficient power detection laser contributes significantly to sensitivity since the more light is sent to the surface, the more is collected. In the case of industrial materials there is essentially no limit on the power or energy that can be sent for detection onto the surface, except in very special cases. Damage in industrial laser-ultrasonics is usually caused by ablation from the generation laser and not the detection laser.
On the other hand, biomedical optical imaging is presently a very active research and application area. Optical diagnostic techniques can be separated into several categories depending upon their sensitivity to optical properties of tissues, namely absorption, scattering or fluorescence and if they are based on the detection of essentially unscattered light (i.e. ballistic photons) or diffused light. Imaging technologies that rely on ballistic or quasi-ballistic (snakelike) photons include conventional microscopy as well as many of its variants such as confocal microscopy, phase contrast microscopy, and multiphoton spectroscopy. Those imaging methods rely on ballistic photons and are essentially limited by the light diffraction limit. Optical coherence tomography (OCT) relies also on the detection of backscattered ballistic or quasi-balistic photons. This technique still provides a high transverse resolution (typically of the order of 1 μm), a good depth resolution (typically 10 μm) but is limited in terms of the probing depth, typically 1 to 2 millimeters in biological tissue. Diffuse optical tomography (DOT) relies instead on the detection of diffuse photons but has a poor spatial resolution (1 cm is typical).
Optical imaging has been recognized to be a powerful biomedical imaging technique, that depends on scattering or absorption within the sample. For example, changes of reflection between tissue layers could indicate the presence of cancer or the accumulation of fat (e.g. underneath an artery wall). Blood vessels are readily detected by the optical absorption of blood and the optical absorption of oxyhemoglobin and deoxyhomoglobin indicate tissue perfusion and metabolism, as well as the presence of cancerous tumors (angiogenesis). Another application of interest is the probing of blood oxygenation in sensitive tissues, such as a retina's blood vessels. Inadequate oxygenation or oxygen consumption within the retina has been linked to many severe eye diseases causing loss of vision.
Biological and medical imaging of tissues also includes purely ultrasonic methods such as ultrasonography (US), which provides a lower resolution (of the order of 1 mm) but with a much higher penetration depth (typically 10 cm) owing to the low scattering of ultrasonic waves in biological tissue (compared to optical waves). Higher transverse resolution can be obtained using higher frequencies but at the cost of the penetration depth. Ultrasonic methods mainly provide information about the mechanical properties of tissue such as the stiffness and the density using reflections due to ultrasonic impedance mismatch between inclusions such as a tumors and the surrounding tissue. More recent innovations in ultrasonic imaging methods include Doppler US and more advanced image reconstruction methods.
Combining the optical contrast provided by optics with the deeper probing capacity of ultrasound has been recognized as desirable, and shown to be applicable in two different ways: using ultrasound-modulated optical tomography (UOT); and photo-acoustic tomography (PAT).
UOT involves ultrasonic waves produced by a conventional focused transducer usually in direct contact with the tissue using water as the coupling medium. The tissue is then illuminated with a single-frequency laser beam and light transmitted through the tissue is collected and analyzed with a large étendue optical system. The ultrasonic waves modulate the phase of light scattered by the tissue. Different implementations are based either on an extremely narrow optical band-pass filter (CFPI or cryogenic atomic filters using spectral hole burning in rare earth ion doped crystals), a photorefractive interferometer (PRI) in quadratic detection, or a CCD camera followed by suitable signal processing (parallel speckle imaging). The light interacting with the ultrasonic wave is tagged by a frequency shift equal to the ultrasonic frequency. Knowing the position of the ultrasound at any time from the known velocity of propagation and the position of the transducer, the tagged photons provide localized information about the optical properties of objects embedded in the biological sample. It has been shown that this method not only provides information on optical absorption but also on optical scattering.
PAT relies on ultrasonic waves generated optically within one or more localized regions (i.e. an optical absorber) inside a highly optically scattering biological tissue. A generation laser emitting short pulses (typical pulse duration of about 10 ns) is used to illuminate the tissue. Photons then propagate within the tissue following highly randomized optical paths, owing to the high concentration of natural optical diffusers (cells) in biological tissue. Energy from these pulses is absorbed by the optical absorbers, with much greater efficiency than the surrounding tissues. The absorbed heat results in thermal expansion of the optical absorber relative to the surrounding tissues. When the optical illumination is very short, the thermal expansion is sufficiently fast to produce ultrasonic waves emanating from the optical absorber. These ultrasonic waves then propagate in the tissue up to its surface, and a single scanning piezoelectric transducer or a matrix of piezoelectric transducers can be used to detect the ultrasonic waves with a given spatial resolution. Transmission of the ultrasonic waves to the scanning transducer or matrix of transducers requires some coupling medium, which is in practice a gel film applied onto the skin of the small animal or human patient, or a water bath in which the animal or human part is immersed. Mathematical techniques are then used to reconstruct the ultrasonic wave distribution at the time of the generation laser pulse illumination. This distribution corresponds to a three-dimensional mapping of the optical absorption of embedded optically absorbing inclusions.
These combinations of optical sensing and ultrasound require contact or fluid coupling of the ultrasound generation (in UOT) or detection (in PAT) device. There are obviously major drawbacks to applying PAT in some applications, such as small animal imaging where the water bath require immersing the animal with an air carrier (aerophore) as shown in FIG. 2. It is also cumbersome in the case of human breast imaging for the detection of cancer which is another application for which PAT has been developed. This coupling requirement also limits the application of PAT in the case of surgery or endoscopic examination. Furthermore non-contact techniques are desired when probing soft tissues such as an eye, especially in highly sensitive regions like the retina, or other layers in the eye. Non-contact detection by optical means in PAT, as done in industrial laser-ultrasonics, would obviously be desirable but is not feasible with prior art knowledge, essentially because of two opposing requirements: the requirement of using a high power detection laser to get sufficient detection sensitivity; and the requirement of using a low power detection laser to avoid damaging the tissue.
Some efforts to apply non-contact PAT to biological phantoms and tissues have been reported in the past, but the reported approaches have failed to be practical or could not produce adequate sensitivity in the case of a tissue that has not been provided a high reflectivity coating. For example “Non-contact detection of laser-induced acoustic waves from buried absorbing objects using a dual-beam common-path interferometer” by Jacques et al. (SPIE proceedings vol. 3224, pp. 307-318 (1998)) and “Optoacoustic tomography using time-resolved interferometric detection of surface displacement” by Payne et al. (Journal of Biomedical Optics, vol. 8, pp. 273-280 (2003)). In these reported works, low power He—Ne lasers were used, and detection was performed from the mirror-like surface of a liquid by using specular reflection, which means that the probing beam had to be normal to the surface. The phantoms actually used were liquid, either completely transparent or seeded with particles to get scattering properties similar to those of tissues for the generation laser beam. The demonstration of the detection of a blood vessel in the forearm of a human volunteer was also disclosed. The forearm was covered with a water layer a few mm thick. The vessel was only about 1 mm deep. In practice, it is desired to be able to detect inclusions (vessels, tumors) much deeper, certainly beyond the limit of OCT and on unprepared surfaces, such as the human skin, the surface of internal organs exposed during surgery or accessible from an endoscopic examination or the shaved skin of a mouse. These previous works do not teach or suggest that it was possible to apply non-contact PAT in such conditions.
So while non-contact PAT has been applied to industrial materials, in particular to the case of a translucent polymer bonded to metal, as reported in the Symposium on Laser Ultrasonics 2010 (see Remote Photoacoustic Imaging for Material Inspection by T. Berer, A. Hochreiner, B. Reitinger, H. Grün and P. Burgholzer, Journal of Physics, conference series 2010), this is because industrial materials have no effective limit on the detection laser power or energy. Existing non-contact PAT techniques do not teach how to apply non-contact PAT on biological tissue in vivo or ex vivo, and does not even suggest that this would be possible.
The conflicting requirements for high power signaling without damaging tissues have so far led to the conclusion that non-contact optical detection of ultrasonic waves on a living organism is not possible, and no such technique or system has been reported. More specifically, to avoid damage to the tissue, energy of the pulsed detection laser, spot size and number of pulses sent onto a given location have to be below the maximum permissible exposure (MPE). This limit for the 1.06 μm wavelength, which is the wavelength of a detection laser used in industrial laser-ultrasonics and for which high power can be obtained, is given (in J/cm2) by the formula 1.1 CA t0.25 where CA=5 and t is the pulse duration (in seconds) or the total duration of multiple pulses if signal averaging is used to increase sensitivity. This formula indicates for typical single pulse duration of 60 μs a value of about 0.48 J/cm2 or 3.7 mJ over a spot of 1 mm in diameter (corresponding to a peak power of 62 W). Considering first only the reflection by the surface and assuming a tissue index of refraction of 1.35, about 2.2% of the incident light is then reflected. Assuming further a collecting aperture 1 cm in diameter located 10 cm away from the surface (collection half angle=0.05 rd) and a Lambertian scattering surface, the collected power is about 3.4 mW. If a confocal Fabry-Perot in transmission mode is used as a demodulator, the detection limit formula for shot noise limited detection is: Ulim,rms=(λ/(4π S)) (2 hv B/(η ID0))1/2, where λ=1.06 μm, hv is the photon energy (1.9×10−19 J at 1.06 μm), S is sensitivity factor (about 1), B is the electronic bandwidth (taken to be 10 MHz), q is the quantum efficiency (about 0.9), and ID0 is the power received by the detector (about ¼ the input power at the entrance of the confocal Fabry-Perot). This formula gives a detection limit equal to about 0.006 nm (rms). The detection limit in pressure is related to the particle displacement by: Plim,rms=ρVac 2π fu Ulim,rms/2 where ρ is the density (about 1000 kg/m3), Vac the acoustic velocity (about 1480 m/s) and fu the ultrasonic frequency (taken to be 5 MHz). The factor 2 comes from the doubling of the displacement by reflection of the ultrasonic wave at the free surface. We find Plim,rms=140 Pa=1.4 mbar. This value is of the order of what could be produced in PAT (see 12.3 in Biomedical Optics, a textbook by Lihong Wang and Hsin-I Wu, John Wiley & Sons 2007). Therefore the sensitivity calculated based only on the reflection of the surface of the tissue, assuming maximum permissible exposure for 1 mm spot size and other design parameters but neglecting all transmission losses through the various optical elements, is barely sufficient. To be able to receive a signal above the noise, several of the parameters used above will have to be modified, for example if one accepts to double the spot size, incident power could be multiplied by 4 and sensitivity doubles but resolution will be diminished. One notes also that sensitivity is far below what is obtained with piezoelectric detection, which is about 0.5 Pa (Oraevsky et al, Proceedings SPIE vol 3601, 1999).
The calculation above considers only the reflection or scattering by the surface of the tissue but it is known that much more light is backscattered by the internal inhomogeneities of the tissue. However only a fraction of this backscattered light is contributing to the phase modulation signal since a scattered photon has to go through the illuminated surface spot where it originates and to exit within a limited solid angle to be collected. Since there is a small probability for these events to occur, the contribution of the internally scattered light will be small. In addition, as explained below, there are other reasons that make the contribution of the internally scattered light small, even by taking into account the increase of light energy density close to the surface of the tissue (of the order of the light penetration depth, i.e. in practice, of the order of 5-10 mm, see for example section 3.6.3 in Biomedical Optics, a textbook by Lihong Wang and Hsin-I Wu, John Wiley & Sons 2007, cited above).
Subsurface phase modulation comes from two effects associated with the ultrasonic wave: the displacement of the scattering centers and the modulation of the index of refraction. The displacement contribution is similar to the surface contribution. If the spatial extent of the ultrasonic wave is longer than the through-depth extent of the effectively collected photon cloud, all the scattering centers have the same motion as the surface, and the contributions of all internally scattered photons is added to that of the surface, thus producing increased sensitivity. If the ultrasonic pulse is generated at one or a few absorbers, e.g. in the case of PAT-like excitation, in general this condition is not met. It would require an exceptionally large absorber (e.g. very large tumor) relative to the detection spot size or some very particular absorber, surface, and internal acoustic structure to generate a relatively uniform ultrasonic wave across the extent of the effectively collected photon cloud. In the more general case of an ultrasonic wave with a spatial extent smaller than the effectively collected photon cloud, phase shift is due to a moving slice of thickness about equal to the size of the inclusion (e.g. tumor). Mathematically the displacement response of the inclusion is convoluted with the effectively collected photon cloud which entails that high frequencies are cut off. The duration of the received signal is then, in this case, equal to the through-depth extent of the effectively collected photons cloud divided by the acoustic velocity. Thus in the cases of interest, when PAT-like excitation is used, the contributions of the scattering center motion are not indicative of the size of the inclusion, which is critical diagnostic information.
Regarding now the index of modulation mechanism, it should be first noted that this effect is opposite to the one produced by the surface motion, since increasing the pressure causes an increase of the refractive index and thus lengthens the path length and gives an additional delay. Due to the bipolar nature of the pressure caused by ultrasonic waves, there is a cancellation effect, unless the ultrasonic wave extent is larger than the through-depth of the effectively collected photon cloud. In this case, since the surface is free, there is reflection of the pressure wave with an opposite polarity, giving an additional cancellation effect.
As a conclusion regarding the contribution of the internally scattered photons, only a small fraction of them is expected to contribute to the signal and their small contribution will be essentially at low frequencies (below 1 MHz) and may not relate to the size of the inclusion. Signals that provide information on the size of an inclusion will be essentially given by light only scattered by the surface of the tissue.
It would therefore appear inevitable that the sensitivity of optical detection of ultrasonic waves in tissues under non-damaging conditions is generally insufficient, and this confirms the belief that non-contact detection and therefore non-contact imaging of biological tissues using PAT-like excitation is expected to be infeasible, regardless of how desirable it would be.